This disclosure relates generally to diagnostic imaging and, more particularly, to improved energy management for a computed tomography (CT) system.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are transmitted to the data processing system for image reconstruction. Imaging data may be obtained using x-rays that are generated at a single polychromatic energy. However, some systems may obtain multi-energy images that provide additional information for generating images.
When the gantry is rotated, it includes a significant amount of rotational kinetic energy that is typically dissipated in a device such as an electrical resistor. That is, the gantry is braked and the rotational energy of the gantry is converted to electrical energy and dissipated as heat in the electrical resistor. In recent years, gantry speeds have increased to provide improved temporal resolution in CT imaging. As such, the amount of dissipated electrical energy that is lost when braking the gantry has increased substantially.
The amount of energy stored in a rotating body is generally a function of its rotational velocity squared. Thus, as gantry speeds have increased over the years from, for example, from 1 to 3 Hz, the amount of energy available and expended in each braking event has significantly increased as well. In addition the amount of rotating mass has increased over this period. Power of the x-ray tube has increased (average and peak power), detector coverage has increased, and the voltage capability of the generator has increased, as examples. These increases have resulted in larger x-ray tubes, detectors, generators, and heat exchangers, as examples. G-loading increases as a function of gantry rotational velocity squared, as well. Thus, the mass of support structure of the rotating portion of the gantry itself has increased over time to maintain acceptably small component deflections during gantry operation.
All of these trends toward larger components, faster gantry speed, and increased mass of the rotating support structure have led to an increased amount of energy to be dissipated during gantry braking events. As a result, the capability of the dissipating resistor itself has increased in size, leading to greater cost as well. Further, it is expected that gantry speeds will only continue to increase to further improve temporal resolution, to perhaps 5 Hz and faster.
As such, these trends in CT lead to an increasing amount of rotational kinetic energy that is lost when the rotating system is braked, and it is therefore increasingly desirable to recover the energy for useful purposes. A gantry typically includes an uninterruptible power supply (UPS) that provides 3-phase power to a 3-phase transformer. Power is split out from the transformer in various sub-circuits to provide power to: a high voltage generator via an AC-DC converter (to power the x-ray tube with DC power); an axial drive and motor for rotating the gantry with 3-phase power (typically 480 VAC 3-phase); and to power other system electronics (typically 120 VAC 3-phase). In such systems, power flows generally only in one direction, and energy recoverable from the CT gantry cannot be used without additional hardware.
One known system for recovering CT rotational energy includes additional AC/DC and DC/AC converters and an additional battery. This known system therefore includes additional hardware and control operation complexity to enable recovery of the gantry energy, which leads to overall system cost and design complexity.
Therefore, it would be desirable to have an overall cost effective method and apparatus to improve energy management in a CT system.